Composite and its use

ABSTRACT

A porous composite suitable for filling a recess or a through-passing hole in an implant. The composite includes particles A prepared from a bioactive material and particles B prepared from a non-bioactive material or weakly bioactive material, which is sintratable with the bioactive material, such that particles A and particles B have been sintered together to a porous composite. Particles A and particles B are essentially homogeneous in size. Also disclosed is an implant which contains a core and the composite.

This application is a U.S. national stage of International applicationPCT/FI98/00331, filed Apr. 15, 1998.

This invention concerns a porous composite as defined in the claim 1.The invention is also concerned with an implant comprising saidcomposite.

GENERAL DEFINITIONS

The definitions below are to be understood herein as follows:

“Biomaterial” means non-living material, which is intended to be used inthe body of a human or an animal. A biomaterial can be 1) inert, 2)bioactive, or 3) capable of bioresorption (solubilizable).

“Inert” means nonreactivity of the respective biomaterial with a tissue.

“A bioactive material” reacts in the physiological conditions within thebody so that the outermost layer of a block manufactured from saidmaterial is converted to form a chemical bond with the surrounding hosttissue.

An “osteoconductive” material means a material which facilitates thegrowth of newly forming bone along its surface but without giving riseto newly forming bone when introduced, for example, in muscle.

An “osteoinductive” material is generally a so called growth factorisolated from the interstitial matter of bone tissue or madesynthetically, which induces the formation of newly forming bone forexample in muscle.

An “implant” is any manufactured device of an artificial material, suchas an artificial joint or a part of it, a screw, a fixation plate or acorresponding orthopedical or odontological device, which is to beintroduced into a tissue.

“Host tissue” or “tissue” means bone tissue or soft tissue into whichfor example an implant has been surgically introduced.

“Micromotion” means microscopic motion (generally below 500 μm) withinthe interfacial region of a surgical implant and the host tissue causedby a dynamic load.

BACKGROUND OF THE INVENTION AND PRIOR ART

The publications, which have been referred to in order to illustrate thebackground of the invention and the prior art, are incorporated in thedescription of the invention below by reference.

Biomaterials and the biological anchoring thereof

Implants for both medical and odontological purposes have already beenmanufactured from various materials for a long time. Various metals,alloys, plastics, ceramic materials, glass ceramic materials and thenewest or biologically active glasses are distinguished from each othernot only by their durability but also by the properties of theinterfacial layer between the implant and the tissue. Inert materials,such as metals and plastics, do not react with a tissue. In this casethere always exists an interfacial layer between the implant and thetissue because the implant and the tissue form two distinct systems.Bioactive materials such as hydroxyapatite, glass ceramics and bioactiveglasses react chemically with the tissue and produce a relatively strongchemical bond in the interface between the implant and the tissue,especially for the bioactive glasses. The implant and the tissue arethus anchored to each other. The rate of healing of the tissue and thepotential chemical fixation to the implant is dependent on the activityof the implant material towards the tissue.

In designing the outermost layer of the implant it has to be consideredthat implants intended for functional activity are subjected to motionunder a load immediately after the surgical operation. This compromisesthe healing and impairs the final result. In addition, the load is notcommunicated to the flexible bone by the structure of a non-elasticimplant but the interfacial region in question is disturbed and theintegration is blocked. Problems are often generated also by the lack ofbone or the unacceptable quality thereof. If for example a dentalimplant is surgically placed into an insufficient or qualitativelyunacceptable bone, the stability in the early phase is not attained andthe surgical operation fails, if any bone is not generated beforehand.Under the functional conditions mentioned above, the undisturbed healingis not achieved with the currently used implants.

Specific Clinical Problems

1. Mechanical micromotions between the implant and the host tissueprevents the fast integration (osseal joining) within 6-12 weeks, inwhich case the device is left without a permanent firm anchorage to thesurrounding tissue. The lack of this anchorage is known to lead toclinical detachment in an early phase (within 1-2 years) or even anumber of years later and to the need of a repeat surgery (1), (2).

2. One approach is to have the surface of the implant made porous forexample by means of a few millimeters deep three-dimensional surfacestructure constructed from microscopic titanium spheres or from titaniumtape. Newly forming bone is expected to grow from the host tissue intothis surface structure. Such a porous biologically inactive surfacestructure gives rise to a microscopic locking structure towards theingrowing newly forming bone but the mechanical properties of thisattachment do not allow a sufficient adaptation under the controlimposed by the load conditions. The optimal anchoring structure betweenthe implant and the host tissue is in a state of a continuousreadaptation to make the strength of the structure to correspond to theload conditions.

3. It has been shown (3) that the attachment of a metallic bone implant(such as an artificial joint) to the host bone can be facilitated by abioactive coating. The material used most often is synthetichydroxyapatite. It has been demonstrated that hydroxyapatite 1)facilitates the mechanical attachment of an implant to the host boneafter it has been attached firmly by means of a surgical operation, 2)diminishes the interference in the integration of the implant to thehost bone caused by the micromotion, and 3) diminishes the retardationof the integration of the implant caused by local lack of bone and bythe lack of contact to the bone implant. Hydroxyapatite is caused toattach to the surface of the implant by using a spraying technique, inwhich case the coating material is applied to the surface mostly onlyfrom the spraying direction. In the biomechanical and biological sense,the most optimal implant surface forms a three-dimensional structure,wherein the interstitial space of the structure forms a growth space toaccommodate the ingrowing bone tissue. In such a case, healing leads tothe formation of a connective locking structure. The growth of a newlyformed tissue is facilitated, if the porous structure is entirely madeof a bioactive material. In such a case the bioactive coating materialforms a three-dimensional osteoconductive surface for the growth ofnewly forming bone. In exceptionally difficult conditions, where thegrowth of host bone is particularly poor for example because of lowquality or small amount of the bone, the growth of the newly formingbone can optionally be improved by combining an osteoinductivecomponent, which directly promotes the generation of bone, to abioactive coating material.

Although a bioactive coating can improve the integration of the implantto the host bone, it must nevertheless be noted that this technique isassociated with many problems. The combination of two materials whichdiffer by their properties (elasticity, thermal expansion), is atechnically demanding task. The coating of a metallic implant with abioactive ceramic material can lead to the early breakdown of thecoating, its fast corrosion, or slow detachment (delamination). This hasshown to be the most common complication in efforts to use bioceramicmaterials, including hydroxyapatite, as a smooth coating material ofmetallic implants (4), (5) (6).

The optimal approach would be a construction which makes use of theadvantages of a bioactive coating material to ensure early ossificationbut in which the possibility has been taken into account that thepermanent integration can be secured by using other constructionalapproaches concerning the surface.

One problem with implants provided with bioactive coatings is also inthat the bioactive surface, which is rather fragile, is damaged rathereasily in the chasing of the implant into the bone.

OBJECTS OF THE INVENTION

One object of the invention is to provide a new composite, which whencombined into the implant secures both rapid ossification and permanentintegration of the implant.

Another object of the invention is to provide an implant, which allowsthe micromotion of the implant and the surrounding tissue (bone) andnevertheless secures rapid growth leading to the integration of theimplant and the bone.

Still another object of the invention is to provide an implant which canbe chased into the bone without a risk of damaging the bioactivestructural component, which promotes the growth of a newly forming bone.

A further object of the invention is to provide an implant wherein thefracturability and the risk of detachment of the bioactive structuralcomponent are smaller than those of the known implants.

SUMMARY OF THE INVENTION

The invention is characterized by the independent claims.

Thus, according to one aspect, the invention concerns a porouscomposite, which is characterized in that is comprises:

particles A manufactured from a bioactive material, and

particles B, which are manufactured from a material which isnon-bioactive or weakly bioactive and which is sintratable to the saidbioactive material, and that the particles A and particles B have beensintered together to form a porous composite.

According to a further aspect, the invention concerns an implant whichis composed of a core and a bioactive structural component which extendsto the surface of the implant. The implant is characterized in that intothe body has been made a recess or a through-passing hole whichcomprises the above mentioned composite according to the invention, saidcomposite forming the surface layer of the implant at the recess or atthe through-passing hole.

BRIEF DESCRIPTION OF THE DRAWINGS

FIGS. 1A-1C show schematically tissue reactions of the compositeaccording to this invention as a function of time,

FIGS. 2A-2B show schematically the behaviour of the continuous and thediscontinuous coating in the bending of the implant framework,

FIGS. 3A-3C show as cross sections the recesses made into the body ofthe implant,

FIG. 4 shows a hip prosthesis, which has three recesses for thecomposite of the invention,

FIG. 5 show as a cross section a recess made into the body of theimplant, said recess being filled with a composite according to theinvention, wherein the composite is comprised of distinct layers,

FIGS. 6A-6F show the use of the composite according to this invention injoining and bone screws,

FIG. 7 shows a light micrograph of glass spheres manufactured by using atorch spraying technique,

FIGS. 8A-8C show X-ray diffractograms of finely-ground glass-basedcones,

FIG. 9 shows a scanning electron micrograph showing bioactive glassspheres sintered together,

FIG. 10 shows implant cones used in tests in vivo, and

FIGS. 11A-11C show push-out or detachment curves of the cones implantedinto the bones in tests in vivo.

PREFERRED EMBODIMEN

TS AND DETAILED DESCRIPTION OF THE INVENTION In the definition of thisinvention, the bioactive material means a material which under thephysiological conditions dissolves at least partly within a few months,most preferably within a few weeks, preferably in about six weeks. Forexample, a bioactive material can be a bioactive glass, a bioactiveceramic material or a bioactive glass ceramic material.

In the definition of this invention the term “non-bioactive or weaklybioactive material” i.e. the material from which the particles B havebeen prepared, means a material which under physiological conditionsdoes not dissolve within the first few months. For example, thismaterial can be a non-bioactive or weakly bioactive glass; a ceramicmaterial, a glass ceramic material or hydroxyapatite. Thus, thismaterial can be any physiologically acceptable material, the bioactivityof which is clearly lower than the material of the particles A and whichadditionally allows the particles A and particles B to be sinteredtogether to form a porous composite. Particularly preferably, thenon-bioactive or weakly bioactive material (the material of theparticles B) begins to dissolve before the bioactive material (thematerial of particles A) has dissolved completely. In this case thesuperimposed formation of a chemical and mechanical bond between thetissue and the implant with respect to each other is best secured.

Preferably, the particles A and the particles B are essentiallyhomogenous in size and approximately of the same size relative to eachother.

Preferably, the diameter of the particles A and the particles B is inthe range of 100-500 μm.

According to a preferred embodiment the particles are spherical, forexample spheres manufactured by a torch spraying process wherein the rawmaterial is glass. In such a case the particles A are made of bioactiveglass and particles B of glass without or almost without bioactivity.

The problem with many traditional bioactive glasses is that they havepoor workability because they easily crystallize. Such bioactive glassescannot be manufactured into spheres.

The international patent application WO 96/21628 (7) describes new typesof bioactive glasses, the working region of which is suited for themanufacture of glass and which thus allow the production of spheres.Typically, these glasses have the following composition:

SiO₂ 53-60% by weight

Na₂O 0-34% by weight

K₂O 1-20% by weight

MgO 0-5% by weight

CaO 5-25% by weight

B₂O₃ 0-4% by weight

P₂O₅ 0.5-6% by weight

provided that

Na₂O+K₂O=16-35% by weight

K₂O+MgO=5-20% by weight

MgO+CaO=10-25% by weight

The above glasses are particularly suitable for use in this invention asthe bioactive glass, i.e. as starting material for the particles A.

Preferably, the ratio of the amounts of particles A and B in thecomposite is adjusted so that the amount of particles A is ⅕ toapproximately {fraction (1/1)} of the total amount of the composite. Aparticularly suitable mixing ratio is one where the amount of particlesA is about ⅓ of the total amount of the composite.

Of course, the composite of this invention can comprise particles ofseveral bioactive materials and/or several non-bioactive materials orweakly bioactive materials.

FIGS. 1A-1C show a tissue reaction of the composite according to thisinvention or a growing locking structure as a function of time. FIG. 1Arepresents a situation immediately following the surgical placement ofthe implant. Immediately next to the surface of the core 11 of theimplant are positioned spheres C, which are composed, for example, ofthe same material as the core 11. A composite layer 10, which iscomposed of bioactive spheres A and spheres B, which are composed ofnon-bioactive or very weakly bioactive material, is left between thecore 11 and the tissue (bone) 12. The spheres A and B are sinteredtogether into a porous composite 10. FIG. 1B, which shows the situationafter about 6-12 weeks, shows that newly formed bone 12 a has grown intothe pores formed by the spheres A and B. Said newly formed bone 12 aforms together with the composite 10 a microscopic locking structurebetween the bone 12 and the core 11. FIG. 1C, which represents thesituation months or years after placing the implant, shows a microscopiclocking structure, wherein newly formed bone 12 a and spheres B arefound. The bioactive spheres A have been completely dissolved.

The series of FIGS. 1A-1C illustrates the formation of a chemical bondand a mechanical bond. Table 1 summarizes the amount of various bondspresent.

TABLE 1 The types of bonds prevailing in FIGS. 1A-1C 1A 1B 1C Chemicalbond Particle A → bone − +++ − Particle B → bone − ± ++ Mechanical bond− ++ +++

FIGS. 2A and 2B show a continuous coating 10 (FIG. 2A) and anon-continuous coating 10 (FIG. 2B) of the core 11, respectively. Thebending of the core 11 in the case of FIG. 2A in the direction of thearrows results in a large ratio between the elongation of the coating 10and the original length. Therefore, there exists the possibility thatthe above mentioned problems might be encountered. In contrast, when thecore of FIG. 2B bends, the ratio between the elongation of the coating10 and the original length is small. Thus the bioactive structuralcomponent functioning as a non-continuous structure retains its positionmuch better.

The implant according to this invention utilizes the principle ofnon-continuous coating. Into the core 11 of the implant are formed oneor more recesses 13 (FIGS. 3-5) or a through-passing hole, and thecomposite according to this invention is applied into such recesses orholes. Thus, the composite will not cover the surface of the core as acontinuous coating. Instead, the composite layer forms a layer 10extending to the surface only at the recess or recesses 13 (or a hole orholes across the structure). FIG. 4 shows a hip prosthesis having threecircular recesses 13 containing the composite according to theinvention. FIGS. 3A-3C show examples of some profiles of the recesses.In FIG. 3A, the edges 13 a and 13 b of the recess are perpendicular tothe surface of the core 13, in FIG. 3B the recess is widening outwardly,and FIG. 3C shows an outwardly closing or locking recess structure. Theprofile of the recess of FIG. 3C is particularly good because it securesthe holding of the composite therein.

FIG. 5 shows an implant according to this invention wherein thecomposite layer 10 is composed of several sublayers 10 a . . . 10 n. Theadvantage with this structure is that the various sublayers can have adistinct mixing ratio between the particles A and B. The mixing ratiosare preferably chosen so as to cause an increasing content of particlesA in the composite from the innermost sublayer 10 a towards the sublayer10 n in contact with the tissue 12. Particularly preferred is thecomposite layer 10 forming a gradient with respect to bioactivity.

Particularly preferred is an implant wherein the amount of the particlesA in the sublayer 10 a of the composite facing the interior of the coreis {fraction (1/10)} of the amount of the sublayer in question, andwherein the sublayer 10 n to come in contact with the tissue is composedexclusively or almost exclusively of particles A.

In the approach of FIG. 5, if desired, inert particles preferably madeof the material of the core can be sintered into the surface of therecess before the formation or the application of the composite into therecess.

According to one embodiment, the implant of this invention can beprepared so that the composite within a recess or in a through-passinghole is formed by applying the particles A and B into the recess, forexample, as a mixture with an organic binding material. Sintering isthen performed wherein the organic binding material is burned. If thecomposite layer is composed of several sublayers, the particles A and Brequired for each sublayer, respectively, are applied separately andsintered.

According to another embodiment the composite can be shaped into a blockof the desired form and size capable of attaching to the recess or thethrough-passing hole in the inplant core. Such a composite block can becomposed of several sublayers, in which case the different sublayershave a different mixing ratio of particles A and B so that the contentof particles A increases from the sublayer facing inwardly into theimplant core of the composite towards the sublayer of the composite incontact with the tissue.

By a proper selection of a narrow fraction and a suitable particle sizeand shape, the void space between the particles can be controlled so asto allow newly forming bone with its blood vessels to penetrate into thestructure. When the ossification proceeds, the spheres prepared from,for example, a bioactive glass, are gradually resorbed. This generatesmore space for the bone, whereby the structure of the bone isstrengthened. Therefore, the amount of the biomaterial is diminished asfunction of time. The diminution can be controlled by a proper selectionof bioactive particles which are variable in their bioactivity and intheir size and shape as well as by changing the mixing ratios of thevarious materials. In order to increase the durability of the finalfixation of the bone, it is possible to use an inactive; porousstructure made of the implant material in the bottom of the recesses. Anessential feature of this surface, sintered for example from spheres, isits three-dimensionality. A conductive and inductive osseal contact isformed quickly. A coating made only by using bioactive glass(enamelling) would result only in the generation of a two-dimensionalreaction surface and the healing would be more difficult.

By virtue of the bioactive material in the recesses an active healingreaction takes place already within a few weeks leading to a maturestage within a few months. This represents a noticeable improvement tothe current situation, in which most of the failures are due to the factthat the fixation of the implant does not occur within the first sixweeks.

The composite within the recesses of the implant is intended to functionas a conductive, and in some applications inductive, surface for therapid growth of newly forming bone and the chemical binding of the hosttissue. The functions of recesses made into the implant core (or intothe through-passing holes) can be summarized as follows:

The first object of the recess is to create for the bone tissue ahealing process which is protected mechanically (from the mechanicalmicromotion). The static and dynamic load towards the implant and theconsequent micromotion is thus not directly directed to the interfacebetween the implant and the host tissue. This mechanically protectedinterface between the implant and the host tissue provides optimalconditions for the ossification and for the formation of a chemicalbonding, in other words, undisturbed conditions are created for a fastintegration of the device into the host bone.

The second object of the recess is to protect the surface materialmechanically during the surgical placement of the implant. The implantcan be affixed tightly to a pre-formed site (press-fit fixation) withoutcausing a direct abrasive force to the bioactive material in the recess.The requirements upon the mechanical structural properties of thematerial can thus be less demanding.

The third object of the recess is to diminish the size of the uniformstructure of the bioactive material. Especially for the enamellizedmaterial, the mechanical integrity is improved with the reduction in thesize of the attachment region. Similarly, the coating of the wholecircumference of the device is avoided, which contributes to theimprovement of the mechanical integration durability of the bioactivematerial. Thus, the susceptibility of fracturing and the risk of loss ofthe bioactive structure is diminished.

The fourth object of the noncontinuous bioactive material placed intothe recess, is to partly counteract the different elastic properties ofthe implant core and the bioactive material. The different elasticitiesof the materials can cause problems, for example, for keeping thebioactive structural component attached in the implant under differentconditions of dynamical load.

The fifth object of the recess is to create a macroscopic surfacestructure for the locking of the newly forming bone, which surfacestructure in itself strengthens the mechanical bonding of the implantdue to the ingrowth of newly forming bone. The inclined lockingstructure (FIG. 3C) provides a macroscopic locking structure between thehost tissue and the device.

A porous surface structure which is unreactive (non-bioactive) with thetissue can be formed on the bottom of the recess. This surface structurefulfils the function of creating, when needed, a microscopicthree-dimensional mechanical locking joint between the device and thebone tissue, based on the growth of the newly forming bone. The objectof this bottom structure is to secure a permanent mechanical bonejunction between the implant and the host tissue in those cases wherethe bioactive coating structure has completely eroded. A second objectof the bottom structure is to cover those bearer regions where the useof a bioactive component is unwanted but which are needed to secure thecircular fixation of the device in the wanted direction.

An example of other applications is a tightening joining screw forvarious orthopedical operations to the bone as shown in FIG. 6A. Theapproach of FIG. 6 is suited particularly for osteoporotic bones. Inthis application the joining screws 14 are fixed into the bone withseparate, for example cone-shaped devices 15 having recesses 13, whichin themselves comprise material causing bioactive ossification. Thebioactive agent (or bioactive agents) are attached to the surface of thedevice according to the method described above. The reference numbers 12and 12′ denote bone and the number 12″ denotes marrow. FIG. 6B shows across section of the conical device 15 along the line A—A of FIG. 6A.FIG. 6C shows a plating operation of a fracture 16 of an osteoporotichollow bone 12 and 12′, wherein the metallic plate has been given thenumber 17.

FIG. 6D shows the fixation of a fracture 16 of the navicular of thewrist by using the tightening joining screw described above.

An another application is, as illustrated in FIG. 6E, an ordinary bonescrew 18, which has recesses 13 made for the bioactive material 10. FIG.6F shows a cross section of a bone screw along the line B—B of FIG. 6E.

EXAMPLES Example 1

Preparation of the Glasses

For the experiments described below, two types of glasses were preparedof which a was bioactive and b was very weakly bioactive. The glasseswere prepared by mixing a paste from PA (pro analys) grade rawmaterials. The raw materials were Na₂CO₃, K₂CO₃, MgO, CaCO₃, CaHPO₄*H₂O,H₃BO₃ and fired SiO₂. The composition of the prepared glasses is givenin Table 2.

TABLE 2 The composition of the prepared glasses (weight-%) Glass Na₂OK₂O MgO CaO P₂O₅ B₂O₃ SiO₂ a 6 12 5 20 4 — 53 b 25, 5 — — 11 2, 5 1, 359, 7

After weighing and mixing, the paste was melted in a platinum crucibleat a temperature of 1360° C. with a melting time of 3 hours. The glassmelt was casted in a graphite mould into blocks which were cooled at520° C. for 30 minutes and subsequently in the oven, which was left tocool after switching the power off. The finished glasses were crushedand melted again in order to homogenize the glass mass. The glasses,which had been re-casted and cooled, were crushed and sieved into the250-297 μm fraction, whereafter the sieved crush was treated with amagnet to remove the small iron particles detached during the crushingoperation.

Example 2

Preparation of Glass Spheres

Using a torch spraying technique, the small glass particles were heatedfor a short time to a sufficient degree to have them melted and becomerounded by virtue of the surface tension. After a quick cooling, theglass spheres were collected into a receptacle.

The torch spraying device used in the experiments comprised of acontainer for the crushed glass, a feeding tube, a common input head forthe gases and crushed glass, and a nozzle. A mixture of acetylene andoxygen was used for heating. The nozzle was Castodyn 8000 nozzle nr. 30,which is intended for ceramic spraying. This nozzle gives a sufficientheat to round even the largest particles. The crushed glass flowed intothe nozzle from the container above the device by its own weight. Aftera suitable mixing ratio has been found, the different quantity of heatrequired for the melting of different fractions can be controlled byadjusting the flow rate of the gases. Smaller particles melt faster thanthe larger ones and thus necessitate passing through the flame at agreater velocity, that is a greater flow rate of the gases. A suitablegas flow for the fraction 250-297 μm was 4 dm³/min for acetylene and 6dm³/min for oxygen. A funnel made of stainless steel with a glasscontainer below was used to collect the glass spheres.

In order to assure a good quality of the glass spheres, sieving (ø250-297 μm), magnet treatment and light microscope checking wereperformed immediately after the preparation. After ultrasonic washing inethanol, the spheres were stored in ethanol in a closed vessel.

FIG. 7 shows a light micrograph of glass spheres (ø 250-297 μm)manufactured by torch spraying technique. The glass speres had beenprepared of bioactive glass (glass a, Table 2).

Example 3

Preparation of Glass-Based Cones

The implants used in the experiments described below were prepared bysintering glass spheres prepared according to the previous example intoporous devices having the shape of a truncated cone. For the preparationof the glass cones, the glass spheres prepared from the glasses a and bof Table 2 were used. Two types of glass cones were prepared, type I andtype II. The first type (I) of glass cones was prepared by sinteringglass spheres which were torch sprayed from the glass a of Table 2. Thesecond type (II) of glass cones was prepared by sintering a mixture ofglass spheres of which ⅓ were glass spheres prepared from the glass a ofTable 2 and ⅔ glass spheres prepared from the glass b of Table 2.

FIGS. 8A and 8B show an X-ray diffractometric analysis of a randomlychosen crushed cone. FIGS. 8A and 8B are X-ray diffractograms of thecone type I and the cone type II, respectively. It can be seen fromthese Figures that the glass has retained its amorphic structure afterthe heating processes associated with the preparation of the cones. FIG.8C shows an X-ray diffractogram of a control cone, wherein the observedpeaks demonstrate the occurrence of crystallization in the glassstructure. The control cone was prepared from glass spheres, for which aconventional bioactive glass, in other words glass without potassium ormagnesium oxide, was used as a raw material.

For the sintering a rectangular mold (50×30×20 mm) was prepared fromgraphite, into which ten 14 mm deep holes were made using a cone-shaped4 mm bit. The holes were filled with the prepared glass microspheres,and the mold with the spheres was heated in a preheated Naber L 49 oven.

Both of the cone types I and II were prepared at the sinteringtemperature of 760° C. The sintering time for the cone type I was 5 min15 s and the sintering time for the cone type II was 3 min 40 s.

The cones which were overshrinked (overmelting) during the heating werediscarded and the accepted cones were checked for the thickness of thenecks between the spheres by using a light microscope. The length of thecones was 14 mm and the ø=2.9 mm and 3.9 mm. The finished cones werewashed in ethanol by using an ultrasonic treatment and stored in ethanolin a closed vessel.

FIG. 9 represents a scanning eletron micrograph, which shows bioactiveglass spheres of type I sintered together ø=250-297 μm.

Example 4

Preparation of Titanium-Based Cones

For comparison, a titanium-based cone type was prepared by sinteringtitanium microspheres. Microspheres prepared from medical grade titaniumby atomizing in a protective argon gas were purchased from Comp Tech,Tampere. The spheres were sieved to a fraction 250-297 μm and washedultrasonically in ethanol. Because titanium reacts very easily withoxygen at higher temperatures, the sintering of titanium must be done ina vacuum oven. For the sintering, molds resembling the ones used inExample 3 were prepared by drilling holes with a 4 mm cone-shaped bitinto a graphite block. The blocks were filled with titanium microspheresand the sintering was performed in a vacuum oven at a temperature of1500° C. and with a sintering time of 2 h 30 min. A successful resultwas checked after the sintering by using a light microscope.

FIG. 10 represents cones used as implants in this study. The cone shownon the right represents the glass cone type I and the cone shown in themiddle the glass cone type II from Example 3. The titanium-based conedescribed above is shown on the left. The spheres have a ø=250-297 μm.

Test Results

The durability of the sintering necks observed in FIG. 9 is influencedessentially by not only the behaviour of the glass in the tissue butalso by the successfulness of the sintering. The sintering result, i.e.the mechanical strength of the matrix, is compromised by the sinteringof more than one type of glass together. This is due to the fact thatdifferent glasses have different coefficients of thermal expansion;during the cooling, microfractures develop in the structure of thematrix. In order to clarify the differences in the mechanical strengthof the different matrices, a mechanical compression test was performedon the cones made of glass spheres.

1) Compression strength of the cones

For the compression test, blocks with dimensions corresponding to thoseof the types I and II of the cones made of glass spheres, respectively,were prepared by sawing off the excessive material from the both ends,in which case the cone block to be tested was 4 mm in length, ø=3.3 and3.4 mm, respectively. The compression strength of the titanium cones wasnot measured, because the strength of the sintered titanium cone wouldhave exceeded the maximum load of the measuring device.

The measuring device was composed of an Alwerton compression device anda recorder. In the device, a downward-moving probe proceeding at aconstant velocity compresses a block on a solid platform. The velocitycan be controlled and the probe measures the load upon the block. Thedevice is connected to a recorder and this is arranged to record themaximum load before the disintegration of the block.

The compression strength of the cones made of glass spheres is shown inthe Table 3.

TABLE 3 Glass cone Number of Compression strength type tests (MPa) I 817, 5 ± 3, 9  II 7 5, 0 ± 1, 0

2) Push-out test of the cones

Into the femur of rabbits (n=8) were implanted cones, which representedthe glass cone types I and II described in Example 3, and thetitanium-based cone type described in Example 4. A similar series ofthree cones were implanted into both femurs, one for histomorphologicaldeterminations and the other for biomechanical determinations. The totalnumber of implants was 3×16=48 cones. After a six week follow-up period,the rabbits were sacrificed, the femurs removed and the force (the pushout force) needed for the detachment of the cone from the bone wasdetermined.

The biomechanical push-out test was performed on the same device as thecompression strength test described above. For the test, the ends of thefemur were cut and the bone was split longitudinally. The excess part ofthe implant inside of the bone was removed and the exterior of the bonewas carefully cleaned. The bone was then placed against a solid support.The support had a central hole with dimensions suitable for fitting theother end of the detaching cone. The device was switched to register themaximal load, the compression rate was 0.5 mm/min. In addition, thedevice was connected to a recorder with a paper velocity of 30 mm/min.

A summary of the results of the push-out test is shown in Table 4.

TABLE 4 The push-out force needed for the implanted cones after a sixweek tissue reaction Push-out force Cone type Number of tests (N) Glasscone type I 8 216, 2 ± 20, 6 Glass cone type 8 293, 3 ± 43, 8 IITitanium-based 8 230, 6 ± 15, 4

Given that the contact surface to the bone was the same for all of thecones (the depth of hard bone=1 mm=the height of the cone) the push-outstrength was calculated by dividing the push-out force by the contactsurface of the cone. The push-out strength is shown in Table 5.

TABLE 5 The push-out strength of the implanted cones after a 6 weektissue reaction Push-out Cone type Number of tests strength (MPa) Glasscone type 8 20, 8 ± 2, 0 I Glass cone type 8 28, 3 ± 4, 2 IITitanium-based 8 22, 2 ± 1, 5 cone

FIG. 11 shows the push-out curves for the different cones, wherein thepush-out force is expressed as a function of dislocation (11A=glass conetype I, 11B=glass cone type II, and 11C=titanium-based). The slopes ofthese curves can be used to calculate the so called push-out stiffness,which is a ratio:

push-out force/dislocation

and which describes the stiffness of the implant core during the pushout test. The push-out stiffnesses for the different cones is given inTable 6. The stiffness of the matrix is directly proportional to theslope of the curve.

TABLE 6 The push-out stiffness of the implanted cones after a 6 weektissue reaction Number of Push-out stiffness Cone type tests (N/mm)Glass cone type I 8 301, 6 ± 150, 6 Glass cone type II 8 214, 5 ± 99, 4 Titanium-based cone 8 277, 4 ± 149, 2

Discussion

Testing of the cones made of glass spheres

The matrix sintered from two different types of spheres allows thecombinatory sintering between the spheres of three different types: a—a,a—b, and b—b. Because the glasses differ by their coefficients ofthermal expansion, the necks a—b are weak or partially broken(tensioned) after cooling. Only the necks between the two similarglasses are strong and these are mainly responsible for the mechanicalstrength of the matrix.

The test results of the compression strengths of the cones made of glassspheres (Table 3) demonstrate that the matrix sintered from microspheresprepared from two different glasses is notably weaker than a matrix ofspheres prepared from a single glass. It can be supposed that all thenecks in a matrix of spheres prepared from a single glass are intactafter cooling. The strength of the matrix sintered from a mixture ofglass spheres (glass cone type II) is improved if the ratio between thea/b spheres is diminished, or the fraction of the spheres prepared fromthe bioactive glass (a) is reduced. In this case the mixture of thespheres is made more homogenous and the number of necks is increased. Inthis study the ratio ⅓ was used.

The behaviour of the implanted cones in vivo

1) Cones sintered from the bioactive glass spheres (glass cone type I)

In connection with the implanting of the cones it was possible toobserve immediately the fast penetration of bone marrow fluid into thecone matrix. The matrix was, due to the capillary force, filledcompletely with tissue fluid and blood, so that there was a plentifulamount of reaction surface between the glass and the tissue/tissuefluid.

The bioactive glass reacts with all of its surface, in which case allthe necks are solubilized with time. This leads gradually to theweakening of the matrix with the onset of breaking of the necks.

In the push-out test, which was made after a six-week tissue reaction,it is observed that the cones sintered from bioactive glass are attachedto the bone rather strongly. In spite of the essentially weaker core,the push-out strength (20.8±2.0 MPa) is approximately of the same orderof magnitude as the push-out strength (16-23 MPa) measured for a conemolded from a bioactive glass in previous studies (8). This can beexplained by the ingrowth of newly forming bone into the matrix.Concurrently with the growth of bone into the cone, the bioactive glassis solubilized and the matrix as a whole is weakened. At the maximalload of the push-out value the considerably weakened necks break nearthe outer edge of the cone and the cone is displaced as the newlyingrown bone yields at the edges of the cone. The abrupt rupture of thebond between the implant and the bone is also illustrated by the socalled push-out stiffness of the cone sintered from bioactive glassspheres (301.6±150.6 N/mm), which is of the same order of magnitude asthe cone sintered from titanium spheres (277.4±149.2 N/mm).

2) Cones sintered from a mixture of glass spheres (glass cone type II)

The tissue reaction of this cone type begins vigorously only on thesurface of the bioactive spheres. In contrast, spheres prepared fromvery weakly bioactive glass (b) react or dissolve very slowly. Newlyforming bone has the opportunity to grow into the pores as induced bythe bioactive component. However, mainly the bioactive componentdissolves from the matrix with time and the spheres prepared from glassb remain as a support for the core.

The push-out test shows clearly the push-out strength (28.3±4.2 MPa) ascompared to the corresponding figure for the cone type I or thetitanium-based cone (approximately 22 MPa). Similarly, the push-outstiffness (214.5±99.4 MPa), which illustrates the stiffness of the core,demonstrates that the core is more flexible by virtue of the remainingmatrix composed of the remaining intact glass type b which providessupport for the core. The stiffness of the cone sintered from themixture of spheres is clearly smaller than the corresponding figure forthe cones sintered from the bioactive glass spheres or the titaniumspheres. The structure of the core is markedly more heterogenous than inthe cones sintered from only one type of glass spheres. The newly formedbone grown into the pores gets support in the push-out process from theremaining matrix composed of the glass type b, in which case the bondingbetween the bone and the cone becomes markedly more durable and moreflexible than the bonding between the cones sintered from only bioactiveglass spheres or titanium spheres and the bone. The large standarddeviation in the push-out strength and in the push-out stiffness isexplained by the different compositions of the individual cones preparedfrom the mixture of spheres (which again results from the fact that themixture of spheres from which the various cones were sintered, was notcompletely homogenous). Thus the durability of the matrix is variable.

3) Cones sintered from titanium microspheres

Titanium is an inert material used widely in surgical implants. Thebonding between the implant and the tissue is good but there is nobonding at the interface. In studies performed previously (8) it hasbeen shown that the push-out strength of a smooth titanium cone(approximately 2 MPa) is considerably inferior when compared with thestrength of a corresponding cone molded from a bioactive glass (16-23MPa). In this study the push-out strength of a cone sintered fromtitanium microspheres (22.2±1.5 MPa) was about ten-fold as compared tothe corresponding strength of the smooth titanium cone measured in theabove-mentioned reference and of the same order of magnitude as thepush-out strength of a cone sintered from bioactive glass spheres. Theincreased strength of a porous titanium cone results from not only thecoarseness of the interface but apparently also from theimplant-supporting influence of the newly formed bone grown into theimplant matrix. At the time of detachment the bone strings at theinterface are broken and the cone is detached. The push-out stiffness(277.4±149.2 N/mm) is of the same order of magnitude than the push-outstiffness of the cones sintered from bioactive glass spheres andmarkedly larger than that of the cones (214.5±99.4 N/mm) prepared fromthe mixture of the spheres. The matrix sintered from titanium spherescan not even be supposed to be flexible.

The composite according to this invention can be used to facilitate thebonding of any orthopedical (medical or veterinary) or odontologicalimplant.

The above mentioned embodiments of this invention represent merelyexamples of the application of the idea of this invention. It is evidentto the one skilled in the art that the various embodiments of thisinvention can be varied within the scope of the following claims.

References

1. Freeman M A R, Plante-Bordeneuve P: “Early migration and late asepticfailure of proximal femoral prostheses”. J Bone Joint Surg 76-B:432-438,1994.

2. Karrholm J et al., “Does early micromotion of femoral stem prosthesesmatter?” 4-7-year stereoradiographic follow-up of 84 cementedprostheses. J Bone Joint Surg 76-B: 912-917, 1994.

3. Jaffle W L, Scott D F: “Total hip arthroplasty withhydroxyapatite-coated prostheses”. J Bone Joint Surg 78-A:1918-1934,1996.

4. Ducheyne P, Cuckler J: “Bioactive prosthetic coatings”. Clin Orthop276:102-114, 1992.

5. Ido K et al., “Cementless total hip replacement. Bioactive glassceramic coating studies in dogs”. Acta Orthop Scand 64:607-612, 1993.

6. Pajamäki J: “Bioactive glass and glass-ceramic interfacial reactionsto bone”. Acta Univ Tamperensis vol 406, 1994.

7. Brink et al., WO 96/21628.

8. Ö. H. Andersson et al., “Evaluation of the acceptance of glass inbone”, J. Mater. Sci.: Mater. in Medicine 3(1992) 145-150.

What is claimed:
 1. A porous composite, which is intended to be filledinto a recess or a through-passing hole of an implant, and whichcomprises: particles A prepared from a bioactive material which willreact in the physiological conditions within the body so that anoutermost layer of a block of said bioactive material forms a chemicalbond with surrounding host tissue, and particles B, which are preparedfrom a non-bioactive material or from a weakly bioactive material whichunder physiological conditions does not dissolve within the first fewmonths, wherein the particles A and the particles B are partially meltedtogether to form a porous composite having a three dimensional structurein which individual particles are connected to at least one adjacentparticle but retain a substantially spherical individual shape,characterized in that the particles A and B are essentially homogeneousin size, and wherein said particles A and particles B are approximatelythe same size compared to one another, and have a diameter of at least250 microns.
 2. The composite according to the claim 1, characterized inthat the diameter of the particles A and B is in the range of 100-500μm.
 3. The composite according to claim 1, characterized in that theparticles A and B are rounded, preferably spherical.
 4. The compositeaccording to claim 1, characterized in that the particles A are composedof bioactive glass and that the particles B are composed of glass whichdoes not have any bioactivity or has a weak bioactivity.
 5. Thecomposite according to claim 4, characterized in that the composition ofthe bioactive glass is as follows: SiO₂ 53-60% by weight Na₂O 0-34% byweight K₂O 1-20% by weight MgO 0-5% by weight CaO 5-25% by weight B₂O₃0-4% by weight P₂O₅ 0.5-6% by weight, provided that Na₂O+K₂O=16-35% byweight K₂O+MgO=5-20% by weight MgO+CaO=10-25% by weight.
 6. Thecomposite according to claim 1, characterized in that the mixing ratioof particles A and B has been chosen so that the amount of particles Avaries from about ⅕ to nearly {fraction (1/1)} of the total amount ofthe composite, preferably about ⅓ of the total amount of the composite.7. An implant comprising a core having a recess or a through-passinghole and a bioactive structural component contained within said recessor through-passing hole and which extends to the surface of the implant,wherein said bioactive structural component comprises a layer of aporous composite which comprises: particles A prepared from a bioactivematerial which will react in the physiological conditions within thebody so that the outermost layer of a block manufactured from saidbioactive material forms a chemical bond with surrounding host tissue,and particles B, which are prepared from a non-bioactive material orfrom a weakly bioactive material which under physiological conditionsdoes not dissolve within the first few months, wherein the particles Aand the particles B are partially melted together to form a porouscomposite having a three dimensional structure in which individualparticles are connected to one at least one adjacent particle but retaina substantially spherical individual shape, characterized in that theparticles A and B are essentially homogeneous in size, and wherein saidparticles A and particles B are approximately the same size compared toeach other, and have a diameter of at least 250 microns.
 8. The implantaccording to claim 7, characterized in that the composite layer as suchis comprised of several sublayers wherein the various sublayers have adistinct mixing ratio between the particles A and B so that the contentof particles A in the composite increases from the sublayer facing theimplant core, towards the sublayer of the composite in contact with thetissue.
 9. The implant according to claim 8, characterized in that theamount of particles A in the sublayer of the composite facing theinterior of the core is {fraction (1/10)} of the amount of the sublayerin question, and that the sublayer to come in contact with the tissue,is composed exclusively or almost exclusively of particles A.
 10. Theimplant according to claim 8, characterized in that the composite withina a recess or in a through-passing hole is manufactured so that theparticles A and B are introduced into the recess or the hole andsubsequently sintered.
 11. The implant according to claim 8,characterized in that the particles A and B required in each of thesublayers are introduced separately into the recess or the hole andsubsequently sintered.
 12. The implant according to claim 8,characterized in that inert particles C, preferably prepared from thematerial of the core, have been sintered to the surface of the recess orthe through-passing hole of the implant core before the formation of orthe addition of the composite into the recess or the hole.
 13. Thecomposite according to claim 1, characterized in that, in the sinteringstep, it has been formed into a block of the desired shape and size,which can be attached into the recess or the through-passing hole madeinto the implant core.
 14. The composite according to claim 13,characterized in that the composite block is comprised of severalsublayers wherein various sublayers have a distinct mixing ratio betweenthe particles A and B so that the content of the particles A in thecomposite increases from the sublayer facing the implant core, towardsthe sublayer of the composite to come in contact with the tissue.